Extended field of view ultrasonic imaging with guided efov scanning

ABSTRACT

An ultrasonic diagnostic imaging system produces an extended field of view (EFOV) image. A 3D imaging probe is moved along the skin of a patient above the anatomy which is to be included in the EFOV image. As the probe is moved, images are acquired from a plurality of differently oriented image planes such as a sagittal plane and a transverse plane. As the probe is moved the image data of successive planes of one of the orientations is compared to estimate the motion of the probe. These motion estimates are used to position a succession of images acquired in one of the orientations accurately with respect to each other in an EFOV display format. The motion estimates are also used to display a graphic on the display screen which indicates the progress of the scan to the user as the probe is being moved. The progress may be indicated in terms of probe velocity, distance traveled, or the path traversed by the moving probe.

This invention relates to medical diagnostic ultrasound systems and, inparticular, to ultrasound systems which perform guided panoramic orextended field of view (EFOV) imaging.

Two-dimensional extended field of view (EFOV) or panoramic ultrasoundimaging is a useful tool for visualizing large or long structures (e.g.,femoral artery, carotid artery), which cannot be entirely viewed inconventional ultrasound imaging. In two-dimensional (2D) panoramicimaging, a large number of 2D images are acquired by manually sweeping aprobe having a one-dimensional ultrasonic transducer array substantiallyalong the plane of the 2D image (the azimuth dimension) as described inU.S. Pat. No. 6,442,289 (Olsson et al.) The acquired overlapping imagesare combined to produce a panoramic image by utilizing probe motionestimates, which are typically measured by assessing registrationbetween consecutive overlapping images. The extra-long images can beadvantageously viewed on a wide aspect ratio display as shown in U.S.Pat. No. 6,516,215 (Roundhill). One limitation of conventional extendedfield of view imaging using a 1-D ultrasonic transducer is that motionis only tracked along one imaging plane, which is assumed to be alignedwith the direction of motion. If the direction of motion is not alignedwith the imaging plane of the transducer, there will be geometricdistortion in the panoramic image and reduced potential for accuratemeasurements.

Another way to image long structures is through free-hand scanning. Infree-hand scanning an ultrasound probe is manually scanned in adirection perpendicular to the plane of the image (i.e., in theelevational dimension) to acquire a series of images from differentplanes which are approximately parallel to each other. These images canbe combined to create a three-dimensional (3D) volume known as afree-hand 3D imaging. Free-hand scanning is described in U.S. Pat. Re.36,564 (Schwartz et al.) Free-hand 3D imaging has the ability to displayanatomical structures from different orientations and planes, instead ofrequiring the clinician to mentally interpret the 3D orientation of thestructure from 2D images. Free-hand 3D scanning can generate volumeswhose size is limited only by the size of accessible acoustic windowsand data storage of the ultrasound system, and so has several clinicaladvantages over conventional 3D ultrasound imaging in which the volumesize is limited by the maximum mechanical or electronic sweep angle ofthe probe.

If precise measurements of distances or volumes are to be made from afree-hand 3D panoramic image, the image acquisition should be calibratedso that the sizes and orientations of structures are geometricallyaccurate. In calibrated 3D panoramic imaging, probe motion tracking andreconstruction are important for producing a calibrated volume. Thetracking of an ultrasound probe provides motion estimates, which aredirectly used to compensate for the probe motion during the sweeping ofthe probe. Reliable 3D volume reconstruction is also critical tominimize image quality loss associated with artifacts. Furthermore, itis also important to provide real-time user feedback for assistingscanning over target structures.

Free-hand acquisition with a 1D array probe has been proposed for 3Dpanoramic imaging in, for instance, U.S. Pat. No. 5,899,861 (Friemel etal.) and U.S. Pat. No. 6,572,549 (Jong et al.). In this approach, probemotion is tracked by evaluating the rate of decorrelation of specklepatterns in sequentially-acquired images from different elevationalplanes. However, it is necessary for the speckle patterns to remainpartially correlated in successive images which may not always bepossible, especially during rapid sweeping of the probe. Also, motionestimates based on speckle de-correlation are not very reliable and arestrongly influenced by artifacts such as those from bright specularreflectors. For more reliable motion tracking, external positioningsensors (e.g., magnetic or optical) can be attached to a 1D array probeas described in U.S. Pat. No. 6,517,491 (Thiele et al.) However thesetracking devices suffer from interference and low sensitivity and canexhibit poor accuracy. The approach also requires that additionalequipment to be attached to both the probe and the system, which isinconvenient.

Recently, mechanical 1D array probes have been superseded by theintroduction of electronically-steered 2D array probes. Atwo-dimensional array transducer can electronically scan a volumetricregion over three dimensions by phased steering of the beams. It is notnecessary to mechanically sweep the probe over the body to acquire a 3Dimage, and there are no moving parts in the probe. A 2D array probe canproduce 3D volumetric images in real time, and can also acquire smallerthree dimensional volumes which are stitched together to produce alarger volume image with coordinated looping display of blood flow asdescribed in U.S. Pat. No. 5,993,390 (Savord et al.) However, 2D arrayprobes have the same limitation as mechanical 1D array probes, which isa field of view limited to the region below the probe.

Yet another approach to 3D imaging are recently proposed I-beam orE-beam probes, which contain a primary imaging array and two or threeadjacent perpendicular tracking arrays as described in U.S. Pat. No.6,102,865 (Hossack et al.) However, this approach is limited in that itcan only estimate the probe motion in prearranged directions set by theorientation of the tracking arrays. Moreover, such systems are expensivesince multiple array must be fabricated in a single probe and operatedsimultaneously. Probe position accuracy is limited by the sizes of thetracking arrays, which are usually much smaller than the imaging array.

In accordance with the principles of the present invention, multipleplanar images are acquired using electronic beam steering from a 2Darray probe. These planar images are used to acquire an extended fieldof view image or volume while simultaneously tracking probe motion inmultiple directions. In a preferred embodiment, free-hand 3D panoramicimages are created and displayed with a 2D array probe which canelectronically steer the beam in any arbitrary image plane. A series ofplanar images comprised of B-mode and/or color Doppler (e.g., velocity,power and/or variance) from a primary plane are acquired during a manualsweep of a 2D array probe over target objects. Ultrasound data from thatplane and, if desired, other planes also, is acquired and used to trackprobe motion by computing registration between consecutive imagesacquired along those planes. Motion estimates from each plane arecombined to find the overall motion vector of the probe. The overallmotion estimate is used to compensate for probe motion in reconstructinga calibrated volume from primary planar images. In accordance with afurther aspect of the present invention, partial volumes are displayedduring the sweep of a 2D array probe for real-time user feedback. Inanother embodiment, one or more planar images are displayed to show theprogress of scanning. In addition, the images from a primary plane canbe displayed along with a 2D panoramic image generated from one or moreof the motion estimation planes. In accordance with another aspect ofthe present invention, an icon is displayed during the sweep of theprobe to inform the clinician as to the speed and/or direction of probemotion.

In the drawings:

FIG. 1 illustrates in block diagram form an ultrasonic diagnosticimaging system constructed in accordance with the principles of thepresent invention.

FIG. 2 illustrates a 2D array transducer being moved over the skin of asubject.

FIG. 3 illustrates the acquisition of ultrasound information in twoplanes as a 2D array transducer is being moved.

FIG. 4 illustrates the acquisition of ultrasound information in threeplanes as a 2D array transducer is being moved.

FIG. 5 illustrates the image acquisition technique of FIG. 4 to image ablood vessel of extended length.

FIGS. 6 a-6 c illustrate the change in position of image structures indifferent planes as the 2D array transducer acquiring the images ismoving.

FIG. 7 is a more detailed block diagram of the motion estimationcomponents of the ultrasound system of FIG. 1.

FIG. 8 is a more detailed block diagram of the EFOV subsystem of FIG. 7.

FIG. 9 is a more detailed block diagram of the 3D volume reconstructorof FIG. 8.

FIGS. 10 a-10 d illustrate the development of a 3D EFOV image using a 2Darray probe in accordance with the present invention.

FIG. 11 illustrates a display produced by an ultrasound system of thepresent invention with probe motion indicators which indicate thescanning speed.

FIG. 12 illustrates another display produced by an ultrasound system ofthe present invention with a probe position indicator which indicatesscanning progress.

Referring first to FIG. 1, an ultrasound system constructed inaccordance with the principles of the present invention is shown inblock diagram form. A probe is coupled to the system which includes atwo-dimensional array transducer 500 and a micro-beamformer 502. Themicro-beamformer contains circuitry which control the signals applied togroups of elements (“patches”) of the array transducer 500 and does someprocessing of the echo signals received by elements of each group.Micro-beamforming in the probe advantageously reduces the number ofconductors in the cable 503 between the probe and the ultrasound systemand is described in U.S. Pat. No. 5,997,479 (Savord et al.) and in U.S.Pat. No. 6,436,048 (Pesque).

The probe is coupled to the scanner 310 of the ultrasound system. Thescanner includes a beamform controller 312 which is responsive to a usercontrol 36 and provides control signals to the microbeamformer 502instructing the probe as to the timing, frequency, direction andfocusing of transmit beams. The beamform controller also control thebeamforming of echo signals received by the scanner by its control ofanalog-to-digital (A/D) converters 316 and a beamformer 116. Echosignals received by the probe are amplified by preamplifier and TGC(time gain control) circuitry 314 in the scanner, then digitized by theA/D converters 316. The digitized echo signals are then formed intofully steered and focused beams by a beamformer 116. The echo signalsare then processed by an image processor 318 which performs digitalfiltering, B mode detection, and Doppler processing, and can alsoperform other signal processing such as harmonic separation, specklereduction through frequency compounding, and other desired imageprocessing.

The echo signals produced by the scanner 310 are coupled to the displaysubsystem 320, which processes the echo signals for display in thedesired image format. The echo signals are processed by an image lineprocessor 322, which is capable of sampling the echo signals, splicingsegments of beams into complete line signals, and averaging line signalsfor signal-to-noise improvement or flow persistence. The image lines arescan converted into the desired image format by a scan converter 324which performs R-theta conversion as is known in the art. The image isthen stored in an image memory 328 from which it can be displayed on adisplay 150. The image in memory is also overlaid with graphics to bedisplayed with the image, which are generated by a graphics generator330 which is responsive to the user control 36. Individual images orimage sequences can be stored in a cine memory 326 during capture ofimage loops or sequences.

For real-time volumetric imaging the display subsystem 320 also includesa 3D image rendering processor in a 3D EFOV subsystem 304 (describedmore fully in FIGS. 8 and 9) which receives image lines from the imageline processor 322 for the rendering of a real-time three dimensionalimage which is coupled to the image memory 328 for display on thedisplay 150.

In accordance with the principles of the present invention the 3D EFOVsubsystem produces images for extended field of view imaging. The EFOVimages can be two dimensional planar images as described in theaforementioned Olsson et al. and Roundhill patents, or can be 3D images.The EFOV images are assembled by estimating the probe motion in a motionestimator 302 using the image data provided by cine memory 326. Themotion estimator can track the movement of the probe along the body of apatient by registering the data of successively acquired images, forexample by using a technique called MSAD block matching as described inU.S. Pat. No. 6,299,579 (Peterson et al.) Other motion estimationtechniques such as non-rigid body registration can be used if desired.The block matching technique computes the displacement betweensuccessively acquired images which are at least partially overlapping.When the displacement is computed of images in different planarorientations, a displacement vector of both magnitude and direction canbe computed in three dimensions by the motion estimator 302. Thedisplacement vector provides position information to the EFOV subsystem304 for relatively positioning images acquired from different planes asthe probe is moving. When the successive images are properly locatedwith respect to each other by the EFOV subsystem, a geometricallyaccurate two or three dimensional EFOV image is produced.

The EFOV imaging technique of the present invention may be more fullyunderstood by referring first to FIG. 2, which is a perspective viewdepicting the motion of a 2D array probe as it acquires an EFOV image.Shown is the 2D array transducer 500 with the surrounding probe case andtransducer stack removed for clarity of illustration. In this drawingthe 2D array transducer 500 is moving along the skin surface 2 of apatient and is acquiring image data as it moves. The 2D array transduceris moving in the direction of the large arrow, which is in the “FRONT”direction and away from the “BACK” direction behind the probe. To eitherside of the direction of motion are the “LEFT” and “RIGHT” directions.

With these directional references in mind, reference is now made to FIG.3, which shows two planes “S” (sagittal) and “T” (transverse) in whichimages are acquired as the probe moves in the direction indicated by thearrow. The two planes are depicted in this example as being rectangularin shape, although in a given embodiment they may have other formatssuch as sector or trapezoidal shapes. The two planes S and T are seenextending in the “DOWN” direction from the 2D array transducer 500.While the 2D array transducer 500 is capable of scanning many more imageplanes and indeed the full volume below the array transducer, in thisembodiment it is only necessary to scan the two S and T planes. The needto scan only two planes means that images can be rapidly acquired inquick succession by alternating or interleaving the scanning of the twoplanes. A high acquisition frame rate means that relatively littlemotion will have occurred between images and there will be significantoverlap in the image data content of successive images, improving theability to find the similarity in the image data and calculate themotion vector, and also improving spatial sampling for laterreconstruction of a calibrated volume. The displacement betweensuccessive images can be accurately estimated and a geometricallyprecise EFOV image assembled by correct location of the images withrespect to each other.

In the example of FIG. 3, two types of EFOV images can be formed. One isformed of the succession of S plane images acquired as the 2D arraytransducer moves in the direction of the arrow. Movement in the BACK toFRONT direction will be reflected in the image content of successivelyacquired S plane images. Movement UP and DOWN will also be reflected inthe image content of the S plane. The displacement between successive Tplane images will reveal movement in the LEFT or RIGHT directions, andalso UP and DOWN. Thus, these displacement estimates are used to producea vector of the relative displacement from one S plane image to thenext. The succession of S plane images are then overlaid in alignmentand stitched together to produce a two dimensional EFOV image with itslongest dimension being in the direction of the arrow. It is importantto note that, unlike conventional EFOV imaging using a 1-D transducer,according to the principles of the present invention it is not necessaryfor the probe motion to be aligned exactly with the S plane. Even if themotion deviates from this plane, this deviation will be tracked by the Tplane so that a well calibrated panoramic image can be generated.

The other EFOV image that can be acquired in the example of FIG. 3 is a3D EFOV image compiled from successively acquired T plane images. Thedisplacement vector between successively acquired T plane images iscalculated from the S and T plane image information as previouslydescribed, then used to properly locate successive T plane images withrespect to each other. As T plane images continue to be acquired in thedirection of the arrow, the thickness of the three dimensional image inthe direction of the arrow grows as more T planes of successivedisplacement are in FRONT of the previously acquired and positionedimages. The 3D EFOV image thus has its longest dimension normal to the Tplane.

It is possible to produce both images at the same time, since planarimages are constantly being acquired for both the 2D EFOV and the 3DEFOV images. For instance, an extended volumetric image can be acquiredof the tissue intercepted by successive T plane images, and an extendedcut plane down the center of this volume can be produced by an EFOV ofsuccessive S planes. The S and T planes do not need to be associatedwith particular dimensions of the physical transducer array 500, but canbe acquired with the S plane oriented along either dimension, with the Tplane perpendicular to this. The S plane does not need to be centeredunder the 2D array as shown in FIG. 3, but can be acquired to eitherside of center or at an angular tilt to form an EFOV image of adifferently oriented cut plane of the 3D volume of the 3D EFOV image.Depending on the choice of apertures for the S and T planes, the T planemay not necessarily be perpendicular to the S plane.

The example of FIG. 3 adds a third acquisition plane, a “C”(cross-sectional) plane as planes parallel to the major plane of the twodimensional array are known. The displacement between successive imagesof the C plane is calculated as described above, and will indicatemotion of the 2D array transducer in both the BACK to FRONT and the LEFTand RIGHT directions. With successive T plane images providingdisplacement information in the UP and DOWN directions, it is seen thata three dimensional displacement vector can be computed without the needfor any information from the S plane. This means that, for a 2D EFOVimage compiled from images from the S plane, the S plane acquisition canbe optimized for imaging and the T and C plane acquisition can beoptimized for displacement measurement. Alternatively, if successive Tplane images are to be assembled in a 3D EFOV image, the C and S planeacquisitions can be optimized for displacement measurement while the Tplane acquisitions are optimized for imaging. The sizes of thedisplacement measurement images can be made smaller or their linespacing increased, for example, while the image plane is kept large andwith high line density for a high resolution image.

It is also possible in the example of FIG. 4 to assemble a 2D EFOV imagefrom successive C plane images, if desired. When a C plane is acquired,the volume data between the C plane and the probe is obtained withoutany additional acoustic transmission. The successive sequences of volumedata can be aligned by 3D registration to find the six degree of freedomof the probe (three translation and three rotation) in the entireacquisition process, which allows accurate reconstruction of either a 3Dor a 2D EFOV image.

It will be appreciated that motion which is normal to a plane will causethe image content of the image plane to rapidly decorrelate from oneplane to another, making displacement estimation in that directionproblematic. However, the use of multiple, differently oriented planesenables such elevational displacement to occur in the plane of anotherimage, where image to image correlation will remain high anddisplacement can still be accurately estimated.

FIG. 5 depicts the EFOV scanning of a blood vessel V in the body by useof the T, S, and C planes. In this example an EFOV image is producedover a considerable length of the blood vessel V, enabling much or allof the characteristics of the vessel to be diagnosed in a single image.As the 2D array transducer probe is moved, the S plane is kept inalignment with the center of the blood vessel V so that a cut planeimage down the vertical center of the vessel can be produced. The depthof the C plane is set so that this plane will continue to intersect thehorizontal center of the blood vessel if the vessel remains at aconstant depth in the body. The T plane continues to intercept the bloodvessel and its surrounding tissue as the probe moves. This arrangementcan be used, for instance, to simultaneously produce multiple EFOVimages, a 3D EFOV image of the blood vessel V from successive T planeimages, and orthogonally oriented 2D cut planes from successive T and Cplane images.

FIGS. 6 a-6 c depict the different in-plane displacements that can bereadily determined from successions of the differently oriented planes.As FIG. 6 a shows, a structure s₁ which appears to the FRONT in one Simage will appear to have moved to the BACK in the next successive Splane image when the probe is moving in the direction of the arrow, asindicated by the small arrow between the successive positions s₁ and s₂of the structure. Displacements of the structure UP and DOWN will alsoreadily be discernible from the relative positions of the structure inthe successive images. FIG. 6B shows that the position of the s₁structure in one T plane image will change to the s₂ position in asuccessive T plane image when the probe moves to the LEFT. UP and DOWNdisplacement can also be readily discerned from successive T planeimages. Successive C plane images will readily reveal FRONT to BACKmovement of the probe, as well as LEFT and RIGHT movement, as shown inFIG. 6 c.

FIG. 7 is a block diagram showing connections between the motionestimator 302 and the 3D EFOV subsystem 304 of FIG. 1. In thisembodiment the cine memory 326 stores received planar images prior totheir use for EFOV images. B mode images are used for motion estimationin this example because B mode images do not exhibit the pulsation orflash artifacts of some Doppler signals. The motion estimator analyzessuccessive B mode images and estimates displacement vectors from imageto image. These motion estimates are forwarded to the 3D EFOV subsystemwhere they are used to relatively align the acquired images. The imageswhich are aligned may be Doppler images, B mode images, or both, asforwarded to the 3D EFOV subsystem from the cine memory. The motionestimator also produces motion information for the graphics generator330 which uses the information for the display of one or more motionindication icons on the image display as described below.

FIG. 8 is a block diagram illustrating further details of a 3D EFOVsubsystem 304 constructed in accordance with the principles of thepresent invention. Motion estimates from the motion estimator 302 arecoupled to a 2D mosaic builder 310, a 3D Doppler volume reconstructor308, and a 3D B mode volume reconstructor 306. B mode images stored incine memory 326 are supplied to the 2D mosaic builder 310, the motionestimator 302, and the 3D B mode volume reconstructor 306. Dopplerimages are supplied to the 3D Doppler volume reconstructor 308. The 3Dvolume reconstructors function in the same manner as the 3D imagerenderer described in U.S. Pat. No. 5,572,291 (Schwartz) in which a 3Dimage is rendered from a plurality of 2D images oriented with respect toeach other as guided by the motion estimates of the motion estimator.One or both 3D renderings, depending upon the image informationsupplied, are coupled to a 3D display subsystem 330 where tissue andflow images of the B mode and Doppler renderings can be merged togetherfor a 3D image of both tissue and flow as described in the '291 patent.Various 3D visualization enhancements can also be applied in the 3Ddisplay subsystem, such as multiplanar reconstruction and surfacerendering. The 3D image is supplied to the image memory 328 for display.The 2D mosaic builder 210 combines partially overlapping 2D image framesto form a 2D EFOV image as described in the aforementioned Olsson et al.patent. The 2D EFOV image is supplied to the image memory 328 also fromwhich it may be displayed independently or together with the 3D EFOVimage of the 3D display subsystem 330. For instance, the 3D displaysubsystem 330 can provide a 3D EFOV image of a volume in the body whilethe 2D mosaic builder provides a 2D EFOV image of a planar slice throughthe 3D volume.

FIG. 9 is a block diagram illustrating further details of a 3D volumereconstructor of an embodiment of the present invention. The illustrated3D volume reconstructor 306,308 can simultaneously form 3D images withboth forward and backward data mapping by means of a forwardreconstructor 332 and a backward reconstructor 334. A 3D volume isreconstructed with forward mapping where input data are directly mappedinto output voxels depending on transformation matrices which areprovided by the motion estimation. While this approach can provide fastresponse with potentially the same resolution as an input 2D image,since new image data is added to the rendered volume as soon as it isreceived, the resulting 3D image may have holes in the reconstructedvolume. On the other hand, the backward mapping method of the backwardreconstructor utilizes inverse transformation matrices so that moredetailed output voxels are generated from input data by means ofinterpolation. With the backward mapping 3D reconstruction, fewer holesappear in the reconstructed volume but the response is slower becausereconstruction must await all of the input data from the scan beforeinter-plane image data can be interpolated. A suitable interpolationmethod which can be used is expressed by

$Q_{0} = {{\frac{d_{1}}{d_{1} + d_{2}}Q_{2}} + {\frac{d_{2}}{d_{1} + d_{2}}Q_{1}}}$

where Q₀ is an interpolated pixel value between two acquired pixels Q₁and Q₂ which is separated from the acquired pixels by distances d₁ andd₂, respectively. Other higher order interpolation techniques such aspolynomial and spline interpolation may alternatively be used. Theillustrated embodiment offers both forward and backward reconstructionso that the user has the choice of a faster or a more detailedreconstruction for display.

FIGS. 10 a-10 d illustrate how a 3D EFOV image appears on the ultrasoundsystem display as a patient is being scanned. As the probe is moved inthe direction indicated by the arrow in FIG. 10 a, image planes whichbisect a blood vessel V are successively added to the front surface ofthe volume 100 being scanned and displayed. As the probe continues itsmovement along the skin of the patient, more and more planes are scannedand added to the front of the volume and the volume grows in thisdimension as shown in the longer EFOV volume 100 of FIG. 10 b. As theprobe moves even further, more planes to the front are added to thevolume, which grows even more as depicted in FIG. 10 c. As previouslyexplained, each new plane is located with respect to the previous planesof the volume by estimating the motion and hence the displacement fromthe last acquired plane to the current one. This displacementinformation may be computed from the image data of one or moreorthogonal scan planes such as image plane C or S as illustrated inFIGS. 3, 4, and 5. Each newly acquired plane is thus added inconsideration of its correct geometrical relationship to the previouslyacquired planes and the volume. The resulting volume is thusgeometrically accurate and capable of supporting quantified measurementsof the structure shown in the volume. FIG. 10 d is an illustration of anactual 3D EFOV image acquired and produced in accordance with thepresent invention. (The black-while grayscale range of pixels of theimage has been reversed from its normal white-on-black in FIG. 10 d forclarity of illustration.)

FIG. 11 illustrates an EFOV display 50 of an ultrasound systemconstructed in accordance with the principles of the present inventionwhich provides guidance to the clinician who is moving the ultrasoundprobe to acquire an EFOV image. In this example a 2D EFOV image 66 iscompiled from a succession of component images acquired in the sagittalplane S of FIG. 3. As the probe moves, component images are continuallyacquired in both the sagittal S and transverse T planes. Images in bothplanes are used by the motion estimator 302 to track the movement of theprobe. Some or all of the transverse and sagittal component images areshown at the top of the display, with 52 being a recently acquiredtransverse plane image and 54 being a recently acquired sagittal planeimage. Graphic icons 56 and 58, produced by the graphics generator 330,are displayed above the component images 52 and 54 and contain a barwhich is highlighted in color to indicate the respective planeorientation. In this example the orientation is indicated as if theplane were being viewed edge-on from above the transducer 500.

In accordance with the principles of the present invention, two probemotion indicators 6 are shown in the center of the display 50. Eachprobe motion indicator has a colored bar 60, 62 which delineates arange, and a small triangle (indicated by arrows) which indicates apoint in the range. In this example each probe motion indicator providesan indication to the clinician of the speed at which the probe ismoving, the two illustrated indicators being for transverse and sagittalmotion, respectively. The position of the graphic triangle above eachbar is computed from the frame to frame displacement estimated by themotion estimator 302 and the known times at which the frames wereacquired. Knowledge of these time and distance values enables a directestimation of velocity in a given direction. In this example, since theclinician is performing EFOV scanning by moving the probe in the FRONTsagittal direction as shown in FIG. 3, the transverse motion indicatortells the clinician whether the motion of his probe is in a straightline along the body or is drifting to the LEFT or to the RIGHT. When thesmall triangle remains centered over the color bar 60, the probe ismoving in a straight line to the FRONT in this example. This means thateach successive sagittal component image used to form the EFOV image 66is in substantially the same image plane. But if the probe begins tomove to the LEFT or the RIGHT, successive component images will not beco-planar and the small triangle will correspondingly move to the leftor the right of the color bar center to indicate this motionaldeviation. The faster the probe moves to the LEFT or the RIGHT, thegreater the displayed deviation of the small triangle from the center ofthe color bar. When the probe resumes movement in a straight line to theFRONT, the small triangle will return to and remain centered again.

The sagittal probe motion indicator including color bar 62 will indicatevelocity in the FRONT direction in this example. Before the scan beginsand the probe is at rest on the patient, the small triangle is above thefar left side of color bar, the zero velocity point. As the clinicianbegins to move the probe and its speed increases, the small trianglebegins to move to the right to indicate the speed of the probe in thesagittal FRONT direction. For a uniform EFOV image it is desirable tomove the probe at a constant speed so that the component images areacquired at substantially evenly spaced temporal and spatial intervals.Using parameters available in a given embodiment such as the desiredlength of the image, the amount of cine memory available, the amount ofcomponent image overlap, and the like, the clinician can set up hissystem so that the small triangle will be centered over the color barwhen the probe is being moved at the desired speed. Thus, to scan astraight expanse of anatomy at the desired speed, the clinician onlyneeds to see that the two small triangles both stay in the center overtheir respective color bars as the probe is being moved. By maintainingthis positioning of the small triangles, the clinician will obtain thehigh quality EFOV image she desires. Since some users may intently watchthe displayed image during acquisition, it is also possible to providean audible alarm, such as bell sound, when the probe is moving too fastor too slow.

It will be appreciated that one or both of the probe motion indicatorscan be denominated in units other than velocity. For example theclinician could decide to acquire an EFOV image over a distance on thebody of 50 cm. The sagittal probe motion indicator bar 62 could then beset to indicate progress toward this distance, starting from zero at theleft side of the color bar 62 and continuing to a covered distance of 50cm at the right. As the clinician commences the scan the small trianglewill continue to move with the displacement of the probe from itsstarting point. When the small triangle has moved all the way to theright end of the color bar, the clinician will know that the 50 cm ofscanning has been completed. Another unit in which the probe motionindicators can be denominated is time, for example.

FIG. 12 illustrates a second EFOV display 50 of an ultrasound systemconstructed in accordance with the principles of the present inventionwhich provides guidance to the clinician who is moving the ultrasoundprobe to acquire an EFOV image. In this example the display does notshow the EFOV image as it is developing, but only shows the componentimages 52 and 54 as they are acquired. It is the component images whichare used by the motion estimator to track the movement of the probe. Asin the previous example the plane orientation icons 56,58 are shownabove the respective component images 52,54.

In the display area above the component images is displayed a graphictrace of the path traversed by the probe as an EFOV scan is performed.In this example the trace 70 is formed by a series of small dots.Orientation arrows below the trace 70 indicate the transverse(LEFT-RIGHT) and sagittal (BACK-FRONT) directions. The trace 70 isproduced from the displacement estimates produced by the motionestimator 302. A series of circles 72 are put down along the trace 70which mark 1 cm increments along the path delineated by the trace.Alternatively the circles 70 could be put down after a given number ofcomponent frames have been acquired or after another increment of timehas transpired during the scan, giving the clinician a sense of theuniformity of the probe motion. At the end of the trace 70 a large dot 8indicates the current probe position during the motion of the probe. Thepath of the trace 70 can be related to a specific point of the arraysuch as by using the axis of intersection of the S and T planes and itspoint of origin on the 2D array as the probe location reference point.

In this example a 3D EFOV image is being compiled from the succession oftransverse planes 52. In this illustrated example the volume of tissuebeing scanned includes a blood vessel V shown in lateral cross-sectionin the transverse component image 52 and in longitudinal cross-sectionin the sagittal component image 54. The clinician can guide her movementof the probe to keep the lateral cross-section of the blood vessel Vcentered in the middle of each successive transverse component image 52as shown in the drawing, and with the succession of sagittal componentimages 54 continuing to show the major longitudinal cut plane of theblood vessel V. The clinician is thereby guided in the acquisition of a3D EFOV image that fully encompasses the blood vessel V. After theguided image acquisition is complete, the clinician can switch to a viewof the 3D EFOV image such as that of FIG. 10 d.

While the preferred embodiment utilizes a 2D array transducer probewhich electronically steers beams over the desired image planes as theprobe is moved, it will be appreciated that the EFOV scan and imageacquisition can also be performed with a mechanical 3D scanning probethat oscillates a 1D array back and forth in the probe as the probe ismoved. For example the 3D mechanical probe shown in US patentpublication 2004/0254466 can be used for EFOV scanning in accordancewith the present invention. Each time the moving 1D array attains agiven position in its oscillatory sweep, an image can be acquired in theimage plane orientation of that array position. One or a few neighboringelements of the array can be continually operated to continually scan anorthogonal plane as the array oscillates. As the probe is moved alongthe skin of the patient these two acquisitions will continually scan twodifferent planes as the probe is moved in an EFOV scan.

1. An ultrasonic diagnostic imaging system which produces a guidedextended field of view (EFOV) image comprising: an ultrasound probeincluding an array transducer which may be moved along the surface of asubject in a front direction; a beamformer coupled to the arraytransducer which controls the probe to repeatedly scan a plurality ofimage planes as the probe is moved along the surface in the frontdirection; an image processor responsive to signals received from thearray transducer to form a sequence of images as the probe is moved; amotion estimator, responsive to a sequence of images which producesestimates of probe motion; an EFOV subsystem responsive to the probemotion estimates which operates to produce an EFOV image from a sequenceof images; a probe motion indicator responsive to the motion estimatorwhich produces an indication of probe motion in the front direction andvariation of probe motion transverse to the front direction; and adisplay responsive to the EFOV subsystem and the probe motion indicatorfor displaying the probe motion indicator as the probe is moved and theEFOV image.
 2. The ultrasonic diagnostic imaging system of claim 1,wherein the probe motion indicator produces an indication of probemotion velocity.
 3. The ultrasonic diagnostic imaging system of claim 2,wherein the probe motion indicator produces an indication of probemotion velocity in the front direction and in a direction lateral to thefront direction.
 4. The ultrasonic diagnostic imaging system of claim 1,wherein the probe motion indicator produces an indication of thedistance of probe travel.
 5. The ultrasonic diagnostic imaging system ofclaim 2, wherein the probe motion indicator comprises a display graphicwhich shows the variance of probe motion from a nominal path of travel.6. The ultrasonic diagnostic imaging system of claim 2, wherein thenominal path of travel comprises the front direction and the directionof variance is lateral of the front direction.
 7. The ultrasonicdiagnostic imaging system of claim 2, wherein the probe motion indicatorcomprises a display graphic which indicates the speed of motion of theprobe relative to a zero velocity.
 8. The ultrasonic diagnostic imagingsystem of claim 1, wherein the probe motion indicator and the EFOV imageare both displayed as the probe is moved.
 9. The ultrasonic diagnosticimaging system of claim 1, wherein the probe motion indicator comprisesa graphic depiction of the path of travel of the probe as it is movedalong the surface.
 10. The ultrasonic diagnostic imaging system of claim9, wherein the graphic depiction of the path of travel includesindicators of increments of the path.
 11. The ultrasonic diagnosticimaging system of claim 2, wherein the indicators are quantifications ofunits of distance or time or number of images acquired as the probe ismoved.
 12. The ultrasonic diagnostic imaging system of claim 2, whereinthe graphic depiction of the path of travel includes a graphic markingthe current position of the probe relative to the path of travel. 13.The ultrasonic diagnostic imaging system of claim 1, wherein the displayfurther comprises a component image display area in which a recentlyacquired image used to produce an EFOV image is displayed.
 14. Theultrasonic diagnostic imaging system of claim 1, wherein the displayfurther comprises a component image display area in which a recentlyacquired image used to produce an estimate of probe motion is displayed.15. The ultrasonic diagnostic imaging system of claim 13, wherein thedisplay further comprises a second component image display area in whicha recently acquired image used to produce an estimate of probe motion isdisplayed, wherein the first and second component images are acquired bythe ultrasound probe from differently oriented image planes.